Immobilized bioactive hydrogel matrices as surface coatings

ABSTRACT

The present invention is directed to a stabilized bioactive hydrogel matrix coating for substrates, such as medical devices. The invention provides a coated substrate comprising a substrate having a surface, and a bioactive hydrogel matrix layer overlying the surface of the medical device, the hydrogel matrix comprising a first high molecular weight component and a second high molecular weight component, the first and second high molecular weight components each being selected from the group consisting of polyglycans and polypeptides, wherein at least one of the first and second high molecular weight components is immobilized (e.g., by covalent cross-linking) to the surface of the substrate.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.10/372,757, filed Feb. 21, 2003, which claims the benefit of priority ofProvisional Application Ser. No. 60/358,625, filed Feb. 21, 2002, bothof which are incorporated herein by reference in their entirety.

FIELD OF THE INVENTION

The present invention relates to cross-linked bioactive hydrogelmatrices that are appropriate for use as immobilized bioactive coatingsto improve the integration and performance of medical devices.

BACKGROUND OF THE INVENTION

The replacement of damaged or diseased tissues or organs by implantationhas been, and continues to be, a long-standing goal of medicine towardswhich tremendous progress has been made. In addition, much progress hasalso been made in the field of treating patients with medical conditionsthrough the implantation of therapeutic medical devices, such as glucosesensors and pacemakers. However, one of the most serious problemsrestricting the use of implants is the wound healing response elicitedby implanted foreign materials (Ratner, B. D., “Reducing capsularthickness and enhancing angiogenesis around implant drug releasesystems” Journal of Controlled Release 78:211-218 (2002)).

Biocompatibility is defined as the appropriate response of the host to aforeign material used for its intended application. Biocompatibilityfurther refers to the interaction between the foreign material and thetissues and physiological systems of the patient treated with theforeign material. Protein binding and subsequent denaturation as well ascell adhesion and activation have been invoked as determinants of amaterial's biocompatibility. Biocompatibility also implies that theimplant avoids detrimental effects from the host's various protectivesystems and remains functional for a significant period of time. Withrespect to medical devices, biocompatibility is determined to a largeextent by the type of acute reaction provoked by implantation. Theextent to which a medical device is integrated with the surroundingtissue depends upon the type of wound healing response that is evoked bythe implanted material. In vitro tests designed to assess cytotoxicityor protein binding are routinely used for the measurement of amaterial's potential biocompatibility. In other words, thebiocompatibility of a material is dependent upon its ability to be fullyintegrated with the surrounding tissue following implantation.

The modulation of this tissue response to an implanted medical devicecomprised of a foreign material is pivotal to successful implantationand performance of such medical devices. Mammalian systems recognizeforeign materials, such as surgically implanted objects or medicaldevices. Upon binding to sites on these foreign materials, a cascade ofevents occur that notify inflammatory cells to surround such materialsand initiate a series of wound healing events which ultimately lead tothe formation of an avascular fibrous capsule surrounding the implanteddevice. The formation of an avascular fibrous capsule can severely limitthe life and usefulness of the implanted medical device, especially insituations where direct contact with specific tissue, such as vasculartissue, muscle tissue, or nerve tissue is vital to the effectiveness ofthe device.

Previous research has shown that the specific interactions between cellsand their surrounding extracellular matrix play an important role in thepromotion and regulation of cellular repair and replacement processes(Hynes, S. O., “Integrins: a family of cell surface receptors” Cell48:549-554 (1987)). Consequently, there has been a heightened interestin work related to biocompatible polymers useful in therapeuticapplications. One particular class of polymers that have proven usefulfor such applications, including contact lens materials, artificialtendons, matrices for tissue engineering, and drug delivery systems, ishydrogels (Wheeler J C, Woods J A, Cox M J, Cantrell R W, Watkins F H,Edlich R F.; Evolution of hydrogel polymers as contact lenses, surfacecoatings, dressings, and drug delivery systems.; J Long Term Eff MedImplants. 1996; 6(3-4):207-17 and Schacht, E., “Hydrogels prepared bycrosslinking of gelatin with dextran dialdehyde” Reactive & FunctionalPolymers 33:109-116 (1997)). Hydrogels are commonly accepted to bematerials consisting of a permanent, three-dimensional network ofhydrophilic polymers with water filling the space between the polymerchains, and they may be obtained by copolymerizing suitable hydrophilicmonomers, by chain extension, and by cross-linking hydrophilicpre-polymers or polymers.

Prior work has shown that a thermoreversible hydrogel matrix, which isliquid near physiologic temperatures, elicits vasculogenesis andmodulates wound healing in dermal ulcers (Usala A L, Dudek R, Lacy S,Olson J, Penland S, Sutton J, Ziats N P, Hill R S: Induction offetal-like wound repair mechanisms in vivo with a novel matrixscaffolding. Diabetes 50 (Supplement 2): A488 (2001); and Usala A L,Klann R, Bradfield J, Ray S, Hill R S, De La Sierra D, Usala M, MetzgerM, Olson G: Rapid Induction of vasculogenesis and wound healing using anovel injectable connective tissue matrix. Diabetes 49 (Supplement 1):A395 (2000)). This bioactive hydrogel material has also been shown toimprove the healing in response to implanted foreign materials;demonstrating a decrease in the surrounding fibrous capsule thicknessand a persistent increase in blood supply immediately adjacent toimplanted materials exposed to this thermoreversible hydrogel (Ravin AG, Olbrich K C, Levin L S, Usala A L, Klitzman B.; Long- and short-termeffects of biological hydrogels on capsule microvascular density aroundimplants in rats. J Biomed Mater Res. 2001 May 1; 58(3):313-8.). Howeverthe use of such a bioactive thermoreversible hydrogel as a biomaterialcoating for a medical device is not practical for devices requiringthree-dimensional or thermal stability. Accordingly, there is a need fora bioactive material that is stable at body temperatures and thusappropriate for use as a coating for use with medical devices,particularly those intended for implantation into mammals.

BRIEF SUMMARY OF THE INVENTION

The invention provides a coated substrate, comprising a substrate havinga surface, and a bioactive hydrogel matrix layer overlying the surfaceof the substrate and immobilized thereon, the hydrogel matrix layercomprising a first high molecular weight component and a second highmolecular weight component, the first and second high molecular weightcomponents each being selected from the group consisting of polyglycansand polypeptides. As used herein, the term “immobilized” refers toaffixation of one or more components of the hydrogel matrix layer viaany chemical or mechanical bonding force or process, such as by covalentattachment. The hydrogel matrix coating may further comprise one or moreenhancing agents selected from the group consisting of polar aminoacids, amino acid analogues, amino acid derivatives, intact collagen,and divalent cation chelators.

Preferred substrates include medical devices. The bioactive hydrogelcompositions are useful both as a layer that serves as a structuralcomponent of a medical device and as a bioactive hydrogel coating thatmodulates the wound healing response to an implanted device and improvestissue integration of a medical device. As a structural component of amedical device, bioactive hydrogel-coated biomaterials can be designedas a space-filling scaffold used to direct tissue organization andvascularization of a medical device. One exemplary use would be as acomposite wound healing device comprising a polymeric microbial barrierand an immobilized bioactive hydrogel of sufficient thickness to providea three dimensional structure to fill anatomic voids such as thoseresulting from donor-site tissue harvesting. As a functional coating ofa medical device, bioactive hydrogel coatings are expected to reduce theavascular capsule surrounding an implanted device, and improve theintimate contact between surrounding tissues and active device elementsand hence the performance of devices such as implanted glucose sensorsfor closed-loop control of diabetes. The compositions are also useful asbioactive hydrogel coatings for artificial organs containing functionaltissue cells, and other passive or active medical devices or implants,and other biosensors.

Also provided is a method of preparing a coated substrate, such as acoated medical device. The method comprises immobilizing a first highmolecular weight component on the surface of the substrate, wherein thefirst molecular weight component is selected from the group consistingof polyglycans and polypeptides. The first high molecular weightcomponent is contacted with a second high molecular weight componentalso selected from the group consisting of polyglycans and polypeptides.The contacting step occurs before, during or after the immobilizingstep. The two high molecular weight components form an immobilizedbioactive hydrogel coating on the surface of the substrate. Preferably,one of the high molecular weight components is a polyglycan, such asdextran, and the other is a polypeptide, such as gelatin.

In a preferred embodiment, the immobilizing step comprises covalentlyattaching at least one of the high molecular weight components to thesurface of the substrate. One or more of the high molecular weightcomponents and/or the surface can be chemically modified, such as byoxidation or amination, to form reactive sites thereon capable ofparticipating in covalent bonding. The high molecular weight componentscan be modified to comprise a plurality of pendant reactive groups alongthe backbone of the molecule or a single reactive group located at eachterminus thereof.

BRIEF DESCRIPTION OF THE DRAWINGS

Having thus described the invention in general terms, reference will nowbe made to the accompanying drawings, which are not necessarily drawn toscale, and wherein:

FIG. 1 illustrates formation of open alpha chains derived from collagenmonomers;

FIG. 2A illustrates the effect of the association of the alpha chainswith dextran;

FIG. 2B illustrates the behavior of the alpha chains without associationof the dextran;

FIG. 3 illustrates the effect of other hydrogel matrix additives;

FIG. 4A illustrates a polyglycan immobilized to a surface of a medicaldevice;

FIG. 4B illustrates a polyglycan immobilized to a surface of a medicaldevice and a polypeptide associated with the polyglycan to form ahydrogel;

FIG. 5 illustrates graphically the effect of a hydrogel matrix inpromoting cell aggregation;

FIG. 6 illustrates graphically the effect of a hydrogel matrix throughinduction of transforming growth factor beta 3;

FIG. 7 illustrates a polyglycan immobilized to a surface of a medicaldevice through a terminal group and a polypeptide associated with thepolyglycan to form a hydrogel;

FIG. 8 illustrates one method of forming an immobilized bioactivehydrogel matrix of the present invention; and

FIG. 9 illustrates a covalently cross-linked hydrogel matrix.

DETAILED DESCRIPTION OF THE INVENTION

The present invention now will be described more fully hereinafter withreference to the accompanying drawings, in which preferred embodimentsof the invention are shown. This invention may, however, be embodied inmany different forms and should not be construed as limited to theembodiments set forth herein; rather, these embodiments are provided sothat this disclosure will be thorough and complete, and will fullyconvey the scope of the invention to those skilled in the art. Likenumbers refer to like elements throughout.

The formulation of a thermoreversible hydrogel matrix providing a cellculture medium and composition for preserving cell viability is taughtby U.S. Pat. No. 6,231,881, herein incorporated by reference in itsentirety. Additionally, a hydrogel matrix useful in promotingvascularization is provided in U.S. Pat. No. 6,261,587, hereinincorporated by reference in its entirety. The thermoreversible hydrogelmatrix taught by these references is a gel at storage temperatures andmolten at physiologic temperatures, and comprises a combination of acollagen-derived component, such as gelatin, a long chain polyglycan,such as dextran, and effective amounts of other components, such aspolar amino acids. The thermoreversible hydrogel matrix taught by thesereferences is discussed below in connection with FIGS. 1-3.

Collagen is a major protein component of the extracellular matrix ofanimals. Collagen is assembled into a complex fibrillar organization.The fibrils are assembled into bundles that form the fibers. The fibrilsare made of five microfibrils placed in a staggered arrangement. Eachmicrofibril is a collection of collagen rods. Each collagen rod is aright-handed triple-helix, each strand being itself a left-handed helix.Collagen fibrils are strengthened by covalent intra- and intermolecularcross-links which make the tissues of mature animals insoluble in coldwater. When suitable treatments are used, collagen rods are extractedand solubilized where they keep their conformation as triple-helices.This is denatured collagen and differs from the native form of collagen,but has not undergone sufficient thermal or chemical treatment to breakthe intramolecular stabilizing covalent bonds found in collagen. Whencollagen solutions are extensively heated, or when the native collagencontaining tissues are subjected to chemical and thermal treatments, thehydrogen and covalent bonds that stabilize the collagen helices arebroken, and the molecules adopt a disordered conformation. By breakingthese hydrogen bonds, the polar amine and carboxylic acid groups are nowavailable for binding to polar groups from other sources or themselves.This material is gelatin and is water-soluble at 40-45° C.

As noted above, gelatin is a form of denatured collagen, and is obtainedby the partial hydrolysis of collagen derived from the skin, whiteconnective tissue, or bones of animals. Gelatin may be derived from anacid-treated precursor or an alkali-treated precursor. Gelatin derivedfrom an acid-treated precursor is known as Type A, and gelatin derivedfrom an alkali-treated precursor is known as Type B. The macromolecularstructural changes associated with collagen degradation are basicallythe same for chemical and partial thermal hydrolysis. In the case ofthermal and acid-catalyzed degradation, hydrolytic cleavage predominateswithin individual collagen chains. In alkaline hydrolysis, cleavage ofinter-and intramolecular cross-links predominates.

FIG. 1 illustrates the hydrolytic cleavage of the tropocollagen 10,forming individual polar alpha chains of gelatin 15. Heatingtropocollagen 10 disrupts the hydrogen bonds that tightly contain thetriple stranded monomers in mature collagen.

FIGS. 2A-2B illustrate stabilization of the matrix monomeric scaffoldingby the introduction of a long-chain polyglycan, such as dextran 20. Asdepicted in FIG. 2A, the dextran 20 serves to hold open the gelatin 15,that has been previously heated, by interfering with the naturalpredisposition of the gelatin 15 to fold upon itself and form hydrogenbonds between its polar groups. In the absence of dextran 20, as shownin FIG. 2B, when the gelatin 15 begins to cool, it will form hydrogenbonds between the amino and carboxylic acid groups within the linearportion of the monomer and fold upon itself, thus limiting availablesites for cellular attachment.

The thermoreversible matrix contains a polyglycan, such as dextran, at atherapeutically effective concentration ranging from, for example, about0.01 to about 10 mM, preferably about 0.01 to about 1 mM, mostpreferably about 0.01 to about 0.1 mM. In one embodiment, dextran ispresent at a concentration of about 0.09 mM.

The thermoreversible matrix also contains gelatin, at a therapeuticallyeffective concentration ranging from, for example, about 0.01 to about40 mM, preferably about 0.05 to about 30 mM, most preferably about 1 to5 mM. Advantageously, the gelatin concentration is approximately 1.6 mM.

In order to increase cell binding, intact collagen may be added in smallamounts to the thermoreversible matrix in order to provide additionalstructure for the cells contained in the matrix. The final concentrationof intact collagen is from about o to about 5 mM, preferably about o toabout 2 mM, most preferably about 0.05 to about 0.5 mM. In oneembodiment, the concentration of intact collagen is about 0.11 mM.

The thermoreversible matrix may additionally contain an effective amountof polar amino acids, which are commonly defined to include tyrosine,cysteine, serine, threonine, asparagine, glutamine, asparatic acid,glutamic acid, arginine, lysine, and histidine. For application in thepresent invention, the amino acids are preferably selected from thegroup consisting of cysteine, arginine, lysine, histidine, glutamicacid, aspartic acid and mixtures thereof, or derivatives or analoguesthereof. By amino acid is intended all naturally occurring alpha aminoacids in both their D and L stereoisomeric forms, and their analoguesand derivatives. An analog is defined as a substitution of an atom orfunctional group in the amino acid with a different atom or functionalgroup that usually has similar properties. A derivative is defined as anamino acid that has another molecule or atom attached to it. Derivativeswould include, for example, acetylation of an amino group, amination ofa carboxyl group, or oxidation of the sulfur residues of two cysteinemolecules to form cystine. The total concentration of all polar aminoacids is generally between about 3 to about 150 mM, preferably about 10to about 65 mM, and more preferably about 15 to about 40 mM.

Advantageously, the added polar amino acids comprise L-cysteine,L-glutamic acid, L-lysine, and L-arginine. The final concentration ofL-glutamic acid is generally about 2 to about 60 mM, preferably about 5to about 40 mM, most preferably about 10 to about 20 mM. In oneembodiment, the concentration of L-glutamic acid is about 15 mM. Thefinal concentration of L-lysine is generally about 0.5 to about 30 mM,preferably about 1 to about 15 mM, most preferably about 1 to about 10mM. In one embodiment, the concentration of L-lysine is about 5.0 mM.The final concentration of L-arginine is generally about 1 to about 40mM, preferably about 1 to about 30 mM, most preferably about 5 to about15 mM. In one embodiment, the final concentration of arginine is about10 mM. The final concentration of L-cysteine, which provides disulfidelinkages, is generally about 5 to about 500 μM, preferably about 10 toabout 100 μM, most preferably about 15 to about 25 μM. In oneembodiment, the final concentration of cysteine is about 20 μM.

The thermoreversible matrix is preferably based upon a physiologicallycompatible buffer, one embodiment being Medium 199, a common nutrientsolution used for in vitro culture of various mammalian cell types(available commercially from Sigma Chemical Company, St. Louis, Mo.),which is further supplemented with additives and additional amounts ofsome medium components, such as supplemental amounts of polar aminoacids as described above.

Advantageously, aminoguanidine may be added to this formulation;however, other L-arginine analogues may also be used in the presentinvention, such as N-monomethyl L-arginine, N-nitro-L-arginine, orD-arginine. The final concentration of aminoguanidine is generally about5 to about 500 μM, preferably about 10 to about 100 μM, most preferablyabout 15 to about 25 μM. In one embodiment, the final concentration isabout 20 μM.

Additionally, the matrix may include one or more divalent cationchelators, which increase the rigidity of the matrix by formingcoordinated complexes with any divalent metal ions present. Theformation of such complexes leads to the increased rigidity of thematrix by removing the inhibition of hydrogen bonding between —NH₂ and—COOH caused by the presence of the divalent metal ions. A preferredexample of a divalent cation chelator that is useful in the presentinvention is ethylenediaminetetraacetic acid (EDTA) or a salt thereof.The concentration range for the divalent cation chelator, such as EDTA,is generally about 0.01 to about 10 mM, preferably 1 to about 8 mM, mostpreferably about 2 to about 6 mM. In a one embodiment, EDTA is presentat a concentration of about 4 mM.

FIG. 3 illustrates the effect of polar amino acids and L-cysteine addedto stabilize the units 25, formed by the gelatin 15 and dextran 20, bylinking the exposed monomer polar sites to, for example, arginine'samine groups or glutamic acid's carboxylic acid groups. Furthermore,disulfide linkages can be formed between L-cysteine molecules (therebyforming cystine), which in turn form hydrogen bonds to the gelatin 15.

The mechanical and thermal characteristics of the thermoreversiblehydrogel described above are to a large extent determined by thethermomechanical properties of one of its major components, gelatin.Gelatin-based matrices typically are molten at near physiologictemperatures and hence cannot be expected to have the requisitedurability and mechanical properties when required for implantation as amedical device in certain applications. Therefore, it is imperative tostabilize these gels through a variety of intermolecular interactionsincluding hydrogen bonding, electrostatic or polar amino acid mediatedbonding, hydrophobic bonding and covalent bonding. Although not wishingto be bound by theory, it is believed that the types of bondingmechanisms described above in association with a polyglycan stabilizepolypeptides such as gelatin. For example, as discussed in more detailbelow, the positively charged polar groups of the collagen-derived alphachains are then able to associate with the negatively charged hydroxylgroups of the repeating glucose units found in, for example, dextran.The gelatin and dextran form a composite bioactive hydrogel containingmacromolecular proteoglycan-type structures.

Unlike the prior art thermoreversible matrix discussed above, thepresent invention provides stabilized compositions comprising animmobilized bioactive matrix that can be used, for example, as a coatingfor implanted medical devices to modulate localized wound healing aroundan implanted medical device, or to produce a localized vasculogenicresponse and encourage tissue integration with the implanted device. Thepresent invention is also directed to a method for manufacturing animmobilized bioactive coating or film of cell scaffolding materialdirectly on a substrate surface, such as the surface of a medicaldevice. The present invention provides a cell attachment scaffold thatsupports the initiation of a series of cell signaling pathways andmodulates the localized wound healing and acute inflammatory cascade inresponse to the implanted foreign material. By “bioactive” is intendedthe ability to facilitate or discourage a cellular or tissue response ofa host to implanted materials. Examples include, but are not limited to,induction of vasculogenesis, inhibition of the formation of a foreignbody response, controlled tissue reorganization around an implantedmaterial or medical device, promotion of cellular adhesion, orregeneration of specific anatomic features such as dermal pegs and reteridges during dermal healing. The term “stabilized” or “stable” isintended to refer to compositions that are water-swellable, poorlysoluble, solid or semi-solid materials at physiological temperature(i.e., about 37° C.) and in physiological fluids (e.g., aqueous bodyfluids having a physiological pH of about 7.4), which remain present inthe host for sufficient time to achieve the intended response.

It is not believed that the immobilized or cross-linked bioactive matrixcoating affects the intrinsic material or chemical properties of theunderlying substrate (e.g., a medical device). Unlike prior art devicesor hydrogels, the present invention is believed to modulate the acuteresponse of a host animal to polymeric materials typically used formedical device manufacture, not by changing the material's properties,but rather by changing the localized tissue response to the implantedmaterial.

The bioactive coatings of the invention can be applied to a surface ofany substrate where such coatings would be useful. In particular,suitable substrates include medical devices. By medical device isintended to include any device, whether active or passive in nature,which may be inserted or implanted into a host organism, such as amammal. The term “medical device” is further intended to encompass anynatural or synthetic device or material, including nucleic acids, whichis used therapeutically either in vivo, such as by implantation into ahuman or animal, or ex vivo to provide therapeutic benefit, whetherintended to be a permanent implant or temporary implant. Such devicesinclude but are not limited to catheters, artificial arteries,artificial organs, medical devices containing cells of either engineeredtissues or isolated tissue fragments or cells derived from naturallyoccurring or genetically engineered sources, ligament replacements, bonereplacements, glucose sensors, coronary pacemakers, lap-bands, monitors,artificial larynxes, prostheses (such as testicular, esophageal,tracheal, and fallopian tube), brain stimulators, bladder pacemakers,bladder stimulators, shunts, stents, tubes, defibrillators,cardioverters, heart valves, joint replacements, fixation devices,ocular implants, cochlear implants, breast implants, neurostimulators,bone growth stimulators, vascular grafts, muscle stimulators, leftventricular assist devices, pressure sensors, vagus nerve stimulators,drug delivery systems, sutures, staples, cell scaffolding materials,active or passive medical devices comprised of gels, pastes or solidsand the like and ex vivo bioreactors for liver, kidney or other organsupport devices. Ex vivo bioreactors are external to the patient's bodyand used temporarily to provide metabolic function pending organtransplantation or other therapeutic intervention. Any foreign objectthat is placed in the body, or in contact with body tissues or fluidswhether for a temporary time period or permanently, may benefit from thepresent invention.

The medical device of the present invention may be rigid or flexible,solid, fibrillar, or woven and may be derived from naturally occurringmaterials or constructed from synthetic materials. Exemplary materialsof construction include acrylates, polyglycolic-polylactic acidcopolymers, polyhydroxybutyrates, polyesters (such as Dacron®), expandedpolytetrafluoroethylene (ePTFE), bioactive glass, ceramics (such ashydroxyapatites), coralline materials, processed tissue (such asdemineralized bone), polycarbonate, polyurethane/polycarbonatecopolymers, metals (such as titanium), and mixtures, composites orsubassemblies thereof. Bioactive glasses generally contain silicondioxide (SiO₂) as a network former and are characterized by theirability to firmly attach to living tissue. Examples of bioactive glassesavailable commercially and their manufacturers include Bioglass®(American Biomaterials Corp., USA, 45% silica, 24% calcium oxide (CaO),24.5% disodium oxide (Na₂O), and 6% pyrophosphate (P₂O₅)), Consil®(Xeipon Ltd., UK), NovaBone® (American Biomaterials Corp.), Biogran®(Orthovita, USA), PerioGlass® (Block Drug Co., USA), and Ceravital® (E.Pfeil & H. Bromer, Germany). Corglaes® (Giltech Ltd., Ayr, UK)represents another family of bioactive glasses containing pyrophosphaterather than silicon dioxide as a network former. These glasses contain42-49 mole % of P₂O₅, the remainder as 10-40 mole % as CaO and Na₂O.

The term “subassemblies” is intended to encompass multiple piececonstruction of the device, wherein the individual pieces of the deviceare constructed of the same or different materials. The term “composite”is intended to encompass devices comprising different active or passivematerials, present to meet specific design requirements for the intendedmedical device.

The present invention provides a stabilized bioactive matrix layer orcoating that overlies an exposed surface of a medical device or othersubstrate and is immobilized thereon. As the present invention is usefulas a coating for any portion of a medical device that may have contactwith body tissues or fluids, either in vivo or ex vivo, both temporarilyand permanently, the term “exposed surface” is intended to encompass anysuch surface of a medical device that is exposed to brief or prolongedcontact with body tissues or fluids. The word “surface” as usedthroughout in reference to a medical device or other substrate istherefore intended to encompass, in particular, any surface of themedical device operatively positioned for exposure to body tissues orfluids.

The matrix layer is formed from at least two high molecular weightcomponents. The high molecular weight components of the bioactivehydrogel matrix are selected from the group consisting of high molecularweight polyglycans, high molecular weight polypeptides, and combinationsthereof. By high molecular weight polyglycan is intended anypolysaccharide consisting of more than about 10 monosaccharide residuesjoined to each other by glycosidic linkages. The polyglycan may consistof the same monosaccharide residues, or various monosaccharide residuesor derivatives of monosaccharide residues. Dextran, a preferredpolysaccharide, typically comprises linear chains of α(1→6)-linkedD-glucose residues, often with α(1→2)- or α(1→3)-branches. Nativedextran, produced by a number of species of bacteria of the familyLactobacilliaceae, is a polydisperse mixture of components.

The polyglycan component preferably has a molecular weight range ofabout 2,000 to about 8,000,000 Da, more preferably about 20,000 to about1,000,000 Da. Unless otherwise noted, molecular weight is expressedherein as number average molecular weight (M_(n)), which is defined as

$\frac{\sum{NiMi}}{\sum{Ni}},$

wherein Ni is the number of polymer molecules (or the number of moles ofthose molecules) having molecular weight Mi.

Any polysaccharide, including glycosaminoglycans (GAGs) orglucosaminoglycans, with suitable viscosity, molecular mass and otherdesirable properties may be utilized in the present invention. Byglycosaminoglycan is intended any glycan (i.e., polysaccharide)comprising an unbranched polysaccharide chain with a repeatingdisaccharide unit, one of which is always an amino sugar. Thesecompounds as a class carry a high negative charge, are stronglyhydrophilic, and are commonly called mucopolysaccharides. This group ofpolysaccharides includes heparin, heparan sulfate, chondroitin sulfate,dermatan sulfate, keratan sulfate, and hyaluronic acid. These GAGs arepredominantly found on cell surfaces and in the extracellular matrix. Byglucosaminoglycan is intended any glycan (i.e. polysaccharide)containing predominantly monosaccharide derivatives in which analcoholic hydroxyl group has been replaced by an amino group or otherfunctional group such as sulfate or phosphate. An example of aglucosaminoglycan is poly-N-acetyl glucosaminoglycan, commonly referredto as chitosan. Exemplary polysaccharides that may be useful in thepresent invention include dextran, heparan, heparin, hyaluronic acid,alginate, agarose, carageenan, amylopectin, amylose, glycogen, starch,cellulose, chitin, chitosan and various sulfated polysaccharides such asheparan sulfate, chondroitin sulfate, dextran sulfate, dermatan sulfate,or keratan sulfate.

By high molecular weight polypeptide is intended any tissue-derived orsynthetically produced polypeptide, such as collagens orcollagen-derived gelatins. Although collagen-derived gelatin is thepreferred high molecular weight polypeptide component, othergelatin-like components characterized by a backbone comprised ofsequences of amino acids having polar groups that are capable ofinteracting with other molecules can be used. For example, keratin,decorin, aggrecan, glycoproteins (including proteoglycans), and the likecould be used to produce the polypeptide component. In one embodiment,the polypeptide component is porcine gelatin from partially hydrolyzedcollagen derived from skin tissue. Polypeptides derived from other typesof tissue could also be used. Examples include, but are not limited to,tissue extracts from arteries, vocal chords, pleura, trachea, bronchi,pulmonary alveolar septa, ligaments, auricular cartilage or abdominalfascia; the reticular network of the liver; the basement membrane of thekidney; or the neurilemma, arachnoid, dura mater or pia mater of thenervous system. Purified polypeptides including, but not limited to,laminin, nidogen, fibulin, and fibrillin or protein mixtures such asthose described by U.S. Pat. No. 6,264,992 and U.S. Pat. No. 4,829,000,extracts from cell culture broth as described by U.S. Pat. No.6,284,284, submucosal tissues such as those described in U.S. Pat. No.6,264,992, or gene products such as described by U.S. Pat. No. 6,303,765may also be used. Another example of a suitable high molecular weightpolypeptide is a fusion protein formed by genetically engineering aknown reactive species onto a protein. The polypeptide componentpreferably has a molecular weight range of about 3,000 to about3,000,000 Da, more preferably about 30,000 to about 300,000 Da.

In a preferred embodiment, gelatin and dextran are components of thebioactive matrix of the present invention. For ease of describing theinvention, the terms “gelatin” and “dextran” are used throughout withthe understanding that various alternatives as described above, such asother polyglycan and polypeptide components readily envisioned by thoseskilled in the art, are contemplated by the present invention.

FIG. 4A illustrates one embodiment of the present invention wherein ahigh molecular weight component of the matrix 20, such as apolysaccharide (e.g., dextran), is immobilized to an exposed surface 30of a medical device 60. In this embodiment, the high molecular weightcomponent 20 is attached to the exposed surface 30 via a plurality ofcovalent linkages 40, such as peptide linkages, between the exposedsurface 30 of the medical device 60 and pendant reactive groups alongthe high molecular weight component chain 20. In this manner, a highmolecular weight component of the matrix, such as either dextran orgelatin, can be covalently attached to an exposed surface 30 of amedical device 60 to form an immobilized coating.

In this particular embodiment, the surface of the medical device mustfirst be activated. Surface activation of synthetic materials is wellknown to those skilled in the art of surface modification. For example,surface activation methods are commonly used for the immobilization ofbiomacromolecules during the preparation of affinity chromatographymedia. Common surface modification techniques are outlined in AffinityChromatography: A Practical Approach, Dean et al., IRL Press, 1985 ISBN0-904147-71-1, which is incorporated by reference in its entirety. Othermethods of preparing synthetic or naturally derived surfaces forsubsequent reaction with macromolecular species in solution arewell-known to those skilled in the art.

In one embodiment, reactive amine groups are formed on the surface. Forexample, perfluorinated poly(ethylene-co-propylene) tape (Teflon®) andpoly(ethylene terephthalate) (PET) sheets can be coated with thin aminepolymer layers deposited from a “monomer” vapor of volatile amines usinga radio-frequency glow discharge. The density of the formed amine layercan be varied by selecting appropriate volatile amines. In oneparticular study, low amine density films were prepared usingn-heptylamine, while high amine density films were prepared usingallylamine (See, Kingshott et al., “Effects of cloud-point grafting,chain length, and density of PEG layers on competitive adsorption ofocular proteins” Biomaterials 23:2043-2056 (2002)). The carboxyl groupsof activated dextran or gelatin react with the available amine groups ofthe surface, to form a Schiff base which can then be further reducedusing either sodium borohydride or sodium cyanoborohydride to formpeptide links. The high molecular weight component is thus immobilizedon the surface of the medical device by covalent linkages therebetween.

The extent and uniformity of surface coverage by the immobilized highmolecular weight components can be varied using reaction parameters wellknown to those skilled in the art. Similarly, by varying theconcentration of reactive species in solution above an activatedsurface, the thickness of the immobilized bioactive hydro gel may becontrolled. For example, a thin uniform bioactive hydrogel coating maybe desirable for the improved long-term function of an implanted glucosebiosensor, where the intended device function requires rapidequilibration between the local tissue environment and the sensorinterface for optimal performance. Such methods are well known to thoseskilled in the art. In another example, relatively thick bioactivehydrogel coatings may be desirable for medical devices requiringextensive tissue integration for optimal performance. Cell scaffolds ortissue bulking devices for sphincter repair and regeneration areexamples of such medical devices which may benefit from a designcomposed of an underlying substrate coated with a bioactive hydrogelcoating to provide both structural and mechanical tissue support whileencouraging tissue integration and localized tissue regeneration. Usingmethods outlined above, one can construct bioactive hydrogel coatingsranging in thickness from about 10⁻⁴ to about 10 cm.

The immobilized dextran or gelatin component may be used as a templateupon which a cell scaffolding material similar to the thermoreversiblehydrogel matrix described above may be constructed. For example, atleast one additional high molecular weight component (e.g., gelatin),and at least one enhancing agent may be added to the immobilized highmolecular weight component (e.g., dextran) to form an immobilizedbioactive hydrogel matrix on the surface of the medical device. Therelative amounts of the various hydrogel ingredients may be varied toobtain a wide range of desirable therapeutic and biomechanicalproperties. In one embodiment, the same concentrations as used in thethermoreversible matrix formulation discussed above are used.

By “enhancing agent” or “stabilizing agent” is intended any compoundadded to the hydrogel matrix, in addition to the two high molecularweight components, that enhances the hydrogel matrix by providingfurther stability or functional advantages. Suitable enhancing agents,which are admixed with the high molecular weight components anddispersed within the hydrogel matrix, include many of the additivesdescribed earlier in connection with the thermoreversible matrixdiscussed above. The enhancing agent can include any compound,especially polar compounds, that, when incorporated into thecross-linked hydrogel matrix, enhances the hydrogel matrix by providingfurther stability or functional advantages.

Preferred enhancing agents for use with the stabilized cross-linkedhydrogel matrix include polar amino acids, amino acid analogues, aminoacid derivatives, intact collagen, and divalent cation chelators, suchas EDTA or salts thereof. Polar amino acids is intended to includetyrosine, cysteine, serine, threonine, asparagine, glutamine, asparticacid, glutamic acid, arginine, lysine, and histidine. The preferredpolar amino acids are L-cysteine, L-glutamic acid, L-lysine, andL-arginine. Suitable concentrations of each particular enhancing agentare the same as noted above in connection with the thermoreversiblehydrogel matrix. Polar amino acids, EDTA, and mixtures thereof, arepreferred enhancing agents. The enhancing agents can be added to thematrix composition before, during, or after immobilization of a highmolecular weight component to the surface of the medical device.

The enhancing agents are particularly important in the stabilizedcross-linked bioactive hydrogel matrix because of the inherentproperties they promote within the matrix. The hydrogel matrix exhibitsan intrinsic bioactivity that will become more evident through theadditional embodiments described hereinafter. It is believed theintrinsic bioactivity is a function of the unique stereochemistry of thecross-linked macromolecules in the presence of the enhancing andstrengthening polar amino acids, as well as other enhancing agents.

For example, aggregation of human fibroblasts exposed to bioactivehydrogels has been observed, while aggregation is not observed whenfibroblasts are exposed to the individual components of the bioactivehydrogel. Results from numerous (over fifty) controlled experiments haveshown that normal neonatal human skin fibroblasts form multi-cellaggregates when exposed to the complete thermoreversible hydrogelformulation at 37° C., while no such cell aggregating activity isdemonstrated using formulations in which the bioactive copolymer is notformed. The aggregated cells form tightly apposed cell clusters withinterdigitating cytoplasmic processes, while cells treated withformulations lacking the copolymer remain round and without surfaceprojections. As shown in FIG. 5, in a sample of human fibroblastsexposed to a bioactive hydrogel comprising dextran and gelatin, at least80% of the cells present were in an aggregated state while less than 20%of the cells present remained as single cells. The opposite effect wasobserved in samples where the human fibroblasts were exposed to collagenmonomer alone, carbohydrate alone, or were left untreated. In samplesexposed to collagen monomer alone, approximately 75% of the cellsremained in a single cell configuration while only about 25% of thecells were in an aggregated state. Nearly the same effect was observedin samples exposed to carbohydrate alone. In samples that were leftuntreated, approximately 60% of the cells remained in a single cellstate while only about 40% of the cells were in an aggregated state.

In a preferred embodiment, dextran is immobilized on an exposed surfaceof the device and gelatin is added, thereby forming a copolymer with thedextran through hydrogen bonding and polar interactions. This embodimentis shown in FIG. 4B where dextran 20, having gelatin 15 associatedtherewith, is immobilized on an exposed surface 30 of a medical device60. These interactions may then be further stabilized through subsequentcovalent bonding mediated by added reactive species (i.e., enhancingagents). The finished product is a stabilized, bioactive hydrogel whichfunctions as a cell attachment scaffold having a localized effect oncellular responses, thereby improving the long-term performance of themedical device.

One such effect on cellular response is illustrated in FIG. 6, whichprovides a graphical representation of the results of one study of geneexpression in normal human neonatal skin fibroblasts. That studydemonstrated a marked induction of transforming growth factor beta 3(TGF-β3) following hydrogel exposure. Expression of this gene isassociated with scarless wound healing as seen during fetal development.Conversely, in the same cells, transforming growth factor beta 1(TGF-β1), which is instrumental in scar formation during adult woundhealing, was not induced by hydrogel exposure reflecting the ability ofthe hydrogel to facilitate a specific character of response in apopulation of tissue cells.

In this embodiment, where dextran is immobilized on the surface andgelatin is added, the dextran, containing predominantly relativelyunreactive hydroxyl groups, requires activation to convert the hydroxylgroups to the more reactive aldehyde groups suitable for cross-linkingto the surface. This must be done prior to contacting the surface of themedical device, which has previously undergone surface modification,such as by the method described above for forming reactive amine groups.For instance, the dextran, or other polyglycan component, can bemodified, such as by oxidation, in order to cross-link with the modifiedsurface of the medical device. One known reaction for oxidizingpolysaccharides is periodate oxidation. The basic reaction processutilizing periodate chemistry is well known and appreciated by thoseskilled in the art. Periodate oxidation is described generally inAffinity Chromatography: A Practical Approach, Dean, et al., IRL Press,1985 ISBN0-904147-71-1. The oxidation of dextran by the use ofperiodate-based chemistry is described in U.S. Pat. Nos. 3,947,352 and6,011,008, which are herein incorporated by reference in their entirety.

In periodate oxidation, hydrophilic matrices may be activated by theoxidation of the vicinal diol groups. With a cellulosic surface, orother polysaccharide surface, this is generally accomplished throughtreatment with an aqueous solution of a salt of periodic acid, such assodium periodate (NaIO₄), which oxidizes the sugar diols to generatereactive aldehyde groups (e.g. dialdehyde residues). This method is arapid, convenient alternative to other known oxidation methods, such asthose using cyanogen bromide. Materials activated by periodate oxidationmay be stored at 4° C. for several days without appreciable loss ofactivity. This method can be used to prepare activated biomaterialsurfaces appropriate for polypeptide binding or to prepare solubleactivated polysaccharides to be bound to surfaces containing primaryamine groups.

Periodate oxidized materials, such as dextran, readily react withmaterials containing amino groups, such as an activated surface of amedical device or a polypeptide, producing a cross-linked materialthrough the formation of Schiff's base links. A Schiff base is a namecommonly used to refer to the imine formed by the reaction of a primaryamine with an aldehyde or ketone. The aldehyde groups formed on thecellulosic surface react with most primary amines between pH values fromabout 4 to about 6. The Schiff's base links form between the dialdehyderesidues on the cellulosic surface and the free amino groups on thepolypeptide or activated surface of the medical device. The cross-linkedproduct may subsequently be stabilized (i.e. formation of stable aminelinkages) by reduction with a borohydride, such as sodium borohydride(NaBH₄) or cyanoborohydride (NaBH₃CN). The residual aldehyde groups maybe consumed with ethanolamine. Other methods known to those skilled inthe art may be utilized to provide reactive groups on one of the highmolecular weight components of the matrix.

The immobilized hydrogel matrix coatings of the present invention arebiomimetic, meaning the coating layer imitates or stimulates abiological process or product. Some biomimetic processes have been inuse for several years, such as the artificial synthesis of vitamins andantibiotics. More recently, additional biomimetic applications have beenproposed, including nanorobot antibodies that seek and destroydisease-causing bacteria, artificial organs, artificial arms, legs,hands, and feet, and various electronic devices. The biomimeticscaffolding materials of the present invention may yield therapeuticallyuseful surface coatings that are stable at about 37° C., or bodytemperature.

Once a high molecular weight component, such as dextran or gelatin, hasbeen covalently cross-linked to the surface of the medical device, thesecond high molecular weight component can be added. The two highmolecular weight components, one being covalently cross-linked to thesurface of the medical device, interact through hydrogen bonding andpolar attractions, thereby forming a stabilized copolymer. Additionally,at least one enhancing agent, as described above, can be added tofurther stabilize the hydrogel matrix.

Dextran or gelatin may be immobilized in a pendant, or chain-like,fashion to the surface of the medical device as shown in FIGS. 4A and4B. This approach may be useful for the development of glucose sensorsor other in-dwelling devices requiring improved soft tissue integrationfor long-term function and biocompatibility. This configuration may alsobe useful for guided tissue growth or as a means of modulating cellgrowth and structure within a three-dimensional tissue engineeredconstruct such as a device intended to function as an artificial liveror kidney. The use of a bio-erodible bulk material allows one tofabricate engineered constructs either for controlled drug delivery withlong-term release of a pharmaceutical agent to an induced vascular bed,or the development of guided tissue growth for bulking applications.

In another embodiment, the dextran 20 (or gelatin component) may beattached to the surface via a peptide link at one terminus of thedextran chain, as shown in FIG. 7. As with the embodiment shown in FIG.4, the surface 30 of the device 60 must be activated using commonsurface modification methods as outlined above. One skilled in the artwould readily understand the parameters necessary for immobilizingdextran at one terminus. (See for example, Larm, a et al., “A NewNon-Thrombogenic Surface Prepared By Selective Covalent Binding OfHeparin Via A Modified Reducing Terminal Residue” Biomat Med Dev ArtifOrgans 11:161-73 (1983)). One skilled in the art would also readilyunderstand the methods for immobilizing other macromolecules such aspolypeptides via one terminus. (See for example, Gregorius, K. et al.,“In Situ Deprotection: A Method For Covalent Immobilization Of PeptidesWith Well-Defined Orientation For Use In Solid Phase Immunoassays SuchAs Enzyme-Linked Immunosorbent Assay” Anal Biochem 29994-91(2001), andOlbrich K. C. et al., “Surfaces Modified With Covalently-ImmobilizedAdhesive Peptides Affect Fibroblast Population Motility” Biomaterials17:144-153 (1996)). Those skilled in the art will further recognize thatsurface activation of substrates can achieve the same end ofimmobilizing one of the two high molecular weight components upon asurface without requiring modification of the native macromolecule. Forexample, surface immobilization of proteins to insoluble PVA substrateshas been described previously (See, Manecke G. and Vogt, H. G., J SolidPhase Biochem 4(233) (1979)).

Dextran immobilization occurring through activated terminal groups forms“end-on” immobilized dextran, or “brush-like surfaces” as shown in FIG.7. Here, the gelatin monomer 15 may then be formed around theimmobilized dextran 20 to produce a biomimetic structure 80 designed tofacilitate the activation of cell signaling pathways similar to thosefound during embryonic development. This configuration may lead to amore hydrogel-like surface and may provide a “softer” surface for guidedcell growth. As with the pendant configuration, both permanent andbio-erodible surfaces may be modified in this manner. The extent ofsurface coverage by dextran is dependent on the molecular weight of thedextran and the extent of dextran branching.

FIG. 8 illustrates yet another embodiment, whereby gelatin isimmobilized to the surface of a medical device. Gelatin with its nativereactive primary amines distributed along the polypeptide backbone canbe immobilized to surfaces containing aldehyde groups. Activatedsurfaces may be formed using radio frequency glow discharge treatment ofpolymeric surfaces in the presence of oxygen or other reactive oxidativespecies. This surface treatment forms aldehydes and other reactivespecies on the surface of the treated material, which can subsequentlyreact with and immobilize gelatin directly. Gelatin may also formlinkages, such as peptide linkages, with the surface of the device in apendant fashion or only at a terminus of the gelatin chain. Gelatin isimmobilized or tethered to a surface of a medical device at roomtemperature as shown in Step 1 of FIG. 8. Next, dextran is added and thetemperature elevated to alter the gelatin quaternary structure todisrupt thermally stabilized intramolecular mediated hydrogen bondingproducing a more open polypeptide conformation as shown in Step 2. ForStep 2, the surface temperature should be elevated to at least about 30to about 90° C., preferably about 40 to about 60° C. Various reactivespecies including, but not limited to, polar amino acids and amino acidderivatives as described earlier are also added and allowed to react atthis step. Decreasing the temperature may further assist intermolecularinteractions so that the dextran begins interacting with the gelatin asshown in Step 3. During Step 3, additional components, such as polaramino acids, can also be added to the matrix. Finally, in Step 4, thesurface is once again at room temperature and a stabilized bioactivematrix coating is formed on the surface. The resulting matrix is atightly bonded gelatin/dextran bioactive hydrogel matrix. Subsequentpost-processing steps may include washing to remove excessive reactivespecies from the stabilized hydrogel. Typically, the temperature of thesurface of the medical device varies from about 20 to about 60° C.during the above-described steps. The pH range for the copolymerformation, as described above, is within the physiologic range,preferably between about 6 and about 8, most preferably between about 7and about 7.6.

Another method useful for immobilizing macromolecules, such as gelatinor dextran, to a surface is by way of a mechanical process. For example,a synthetic thermoplastic polymer may be partially swollen in thepresence of a fluid containing a polymer solvent dispersed in water. Theaddition of macromolecules to this fluid allows the added solutes tobecome entrapped within the open, swollen surface of the polymer. Byrapidly exchanging the fluid phase surrounding the swollen polymer, thepolymer de-swells, entrapping the added macromolecular solutes withinthe surface of the polymer.

In yet another embodiment, bioactive hydrogels may be formed directlyusing electrooxidation. In this method, a molten thermoreversiblehydrogel is placed in an electrolytic cell containing two conductiveelectrodes. A potential difference is applied between the electrodes,and oxidizable species in solution (i.e. functional groups such ashydroxyls, and amines) are directly oxidized at the anode. The resultingreactive oxidized compounds condense at the anodic surface to form awater-insoluble hydrogel coating. In this manner, for example, titaniummesh commonly used for craniofacial reconstructive surgery can be coatedwith an immobilized bioactive hydrogel to direct tissue organization andvascularization at the site of the implant.

Additional methods for immobilizing macromolecules are known in the artand include the methods described in the following references, each ofwhich is incorporated by reference in its entirety: (i) Puleo D. A. etal., “A technique to immobilize bioactive proteins, including bonemorphogenetic protein-4 (BMP-4), on titanium alloy” Biomaterials,23:2079-2087 (2002); (ii) Kong, U. et al., “Durable Surface Modificationof Poly(tetrafluoroethylene) by Low Pressure H₂O Plasma TreatmentFollowed by Acrylic Acid Graft Polymerization” Coll Surf B: Biointerface24:63-71 (2002); (iii) Chandy, T. et al., “Use of Plasma Glow forSurface Engineering Biomolecules to Enhance Blood Compatibility ofDacron and PTFE Vascular Prostheses” Biomaterials 21:699-712 (2000);(iv) Bos, G. W. et al., “Proliferation of Endothelial Cells onSurface-Immobilized Albumin-Heparin Conjugate Loaded with BasicFibroblast Growth Factor” J Biomed Mater Res 44:330-340 (1999); (v)Ayhan F. et al., “Optimization of Urerase[sic] Immobilization ontoNon-Porous HEMA Incorporated Poly(EGDMA) Microbeads and estimation ofkinetic parameters” Biores Technol 81:131-40 (2002); (vi) Massia S. P.et al., “Surface Immobilized Dextran Limits Cell Adhesion and Spreading”Biomaterials 21:2253-2261 (2000); (vii) Barie, N. et al., “CovalentPhoto-Linker Mediated Immobilization Of An Intermediate Dextran Layer ToPolymer-Coated Surfaces For Biosensing Applications” Bios Bioelect13:855-860 (1998); (viii) Chevolot, Y., et al., “Immobilization OnPolysytrene Of Diazirine Derivatives Of Mono- And Disaccharides:Biological Activities Of Modified Surfaces” Bioorganic & Med Chem9:2943-53 (2001); (ix) Tsai, C. C. et al., “Effects Of HeparinImmobilization On The Surface Characteristics Of A Biological TissueFixed With A Naturally Occurring Crosslinking Agent (Genipin) An InVitro Study” Biomaterials 22:523-33 (2001); (x) Ito, Y., “MicropatternImmobilization Of Polysaccharide” J Bioinorg Chem 79:88-81 (2000); (xi)Massia, S. P. et al., “Immobilized rgd Peptides On Surface-GraftedDextran Promote Biospecific Cell Attachment” J Biomed Mater Res56:390-399 (2001); and (xii) Dai L., et al., “Biomedical Coatings ByCovalent Immobilization Of Polysaccharides Onto Gas-Plasma-ActivatedPolymer Surfaces” Surf Interface Anal 29:46-55 (2000).

In yet another embodiment of the present invention, the two highmolecular weight components of the hydrogel matrix surface coating maybe cross-linked. As when cross-linking the dextran, or anotherpolyglycan, to the modified surface of the medical device, the dextranmust also first be modified in order to cross-link with the gelatincomponent. For example, partial oxidation of dextran using sodiummeta-periodate produces a polyaldehyde dextran that can be immobilizedon amine derivatized surfaces. Dextran immobilization in the presence ofsodium cyanoborohydride catalytically reduces the formed Schiff base tothe more stable amide covalent bond. The immobilized dextran coatedsubstrate can then be washed to remove the excess reagents, and treatedwith sodium meta-periodate to form additional aldehydes. This tetheredpolyaldehyde dextran can then cross-link with another high molecularweight component, such as gelatin.

The presence of cross-linking between the two high molecular weightcomponents of the hydrogel coating is illustrated in FIG. 9. As shown,in addition to being covalently attached to the exposed surface of themedical device, the dextran 20 can be covalently crosslinked to gelatin15 by linkages 70, thereby forming a crosslinked network 50. Thelinkages 70 either result from reaction of functional groups on thegelatin 15 with functional groups on the dextran 20, or result fromreaction of a bifunctional crosslinker molecule with both the dextran 20and gelatin 15. One method of crosslinking gelatin and dextran is tomodify the dextran molecules 20, such as by oxidation, in order to formfunctional groups suitable for covalent attachment to the gelatin 15.Dextran is modified, such as by oxidation, and stabilized via covalentbonding to gelatin 15, thereby forming a cross-linked network 50.

As noted above, periodate oxidation is one example of a known reactionfor oxidizing polysaccharides that can also be used in this embodimentof the present invention in addition to other embodiments describedpreviously. The reaction scheme can be carried out as before, oxidizingthe sugar diols of the polyglycan, thereby forming reactive aldehydegroups. In this embodiment, the Schiff's base links form between thereactive aldehyde groups and the free amino groups on the polypeptidecomponent of the hydrogel matrix. The cross-linked product may then bestabilized (i.e. formation of stable amine linkages) by reduction with aborohydride, such as sodium borohydride (NaBH₄) or cyanoborohydride(NaBH₃CN), and the residual aldehyde groups may be consumed withethanolamine.

As an alternate method for forming the cross-linked hydrogel coating, amultifunctional cross-linking agent may be utilized as a reactive moietythat covalently links the gelatin and dextran chains. Such bifunctionalcross-linking agents may include glutaraldehyde, epoxides (e.g.bis-oxiranes), oxidized dextran, p-azidobenzoyl hydrazide,N-[α-maleimidoacetoxy]succinimide ester, p-azidophenyl glyoxalmonohydrate, Bis-[β-(4-azidosalicylamido)ethyl]disulfide (BASED),bis[sulfosuccinimidyl]suberate, dithiobis[succinimidyl propionate,disuccinimidyl suberate, 1-ethyl-3-[3-dimethylaminopropyl]carbodiimidehydrochloride, ethoxylated (20) trimethylpropane triacrylate, and otherbifunctional cross-linking reagents known to those skilled in the art.

In one embodiment, 1.5 mL of a 0.5 mg/mL solution ofBis-[β-(4-azidosalicylamido)ethyl]disulfide (BASED) in dimethylsulfoxide (DMSO), is added to a foil-wrapped vessel containing 15 mL ofliquid thermoreversible hydrogel as described above. Photoactivatednon-specific cross-linking of the thermoreversible hydrogel occurs uponexposure of the reactive mixture to long-wavelength light, such as thatprovided by continuous exposure to a 550 watt bulb (flood light used inphotography). Longer exposure times demonstrated better cross-linking.

In another embodiment utilizing a cross-linking agent, polyacrylatedmaterials, such as ethoxylated (20) trimethylpropane triacrylate, may beused as a non-specific photo-activated cross-linking agent. Componentsof an exemplary reaction mixture would include thermoreversible hydrogelheld at 39° C., polyacrylate monomers, such as ethoxylated (20)trimethylpropane triacrylate, a photo-initiator, such as eosin Y,catalytic agents, such as 1-vinyl-2-pyrrolidinone, and triethanolamine.Continuous exposure of this reactive mixture to long-wavelength light(>498 nm) would produce a cross-linked hydrogel network.

In a further embodiment of the present invention, both high molecularweight components are cross-linked to the surface of the medical device.In this embodiment, the surface of the medical device must be activatedprior to contacting the high molecular weight components. For example,bis-oxiranes, such as 1,4-butanediol diglycidoxy ether react readilywith hydroxy- or amino-containing biomaterials at alkaline pH to yieldderivatives which possess a long-chain hydrophilic, reactive oxirane(epoxide), which, in turn, can be reacted with amines, hydroxyls andother nucleophiles. Oxirane-coupled ligands are widely used andextremely stable and the use of a long chain bis-oxirane reagentintroduces a long hydrophilic spacer molecule between the immobilizedhydrogel components and the biomaterial surface which may be desirablein certain applications.

In another embodiment, the high molecular weight components may bemodified in order to form reactive groups capable of reacting with thereactive groups of the activated surface of the medical device prior tocontacting the surface of the medical device. This embodiment is notrestricted by the order in which the high molecular weight componentsare contacted with the surface of the medical device. In one preferredembodiment, the surface of the medical device is activated, such as byradio frequency glow discharge in the presence of amine containingvapors to form reactive amine groups thereon, and modified dextran, suchas oxidized dextran, is added. The dextran is covalently cross-linked tothe surface of the medical device and gelatin is added. The gelatin isthen also covalently cross-linked to the surface of the medical devicethrough the polyaledehyde groups of the tethered dextran.

In another embodiment of the present invention, the surface of themedical device and the two high molecular weight components are allcross-linked to each other, wherein the polyglycan is covalentlycross-linked to the surface of the medical device, the polypeptide iscovalently cross-linked to the surface of the medical device, and thepolyglycan and the polypeptide are cross-linked to each other. Variousmethods for carrying out this embodiment for coating a medical devicewould be envisioned by one skilled in the art. One possible method wouldcomprise coating a polymeric medical device through radiation orelectron beam grafting (See, Muzykewicz K. J. et al., “Platelet adhesionand contact activation time tests on HEMA coated cellulose acetatemembranes” J Biomed Mater Res. 9(5):487-99 (1975) and Venkataraman S. etal., “The reactivity of alpha-chymotrypsin immobilized onradiation-grafted hydrogel surfaces” J Biomed Mater Res. 11(1):111-23(1977)).

The stabilized cross-linked bioactive hydrogel can be used to encouragesite-specific tissue regeneration, including vasculogenesis, in the areasurrounding an implanted medical device with the stabilized cross-linkedbioactive hydrogel immobilized thereon. It is known in the art to useintact collagen, gelatin, or dextran as a carrier to hold and delivergrowth factors and the like in methods designed to promote tissuegrowth. (See, for example, Kawai, K. et al., “Accelerated tissueRegeneration Through Incorporation of Basic Fibroblast GrowthFactor-Impregnated Gelatin Microspheres into Artificial Dermis”Biomaterials 21:489-499 (2000); and Wissink, M. J. B. et al., “Bindingand Release of Basic Fibroblast Growth Factor from Heparinized CollagenMatrices” Biomaterials 22:2291-2299 (2001)). By contrast, the intrinsicactivity of the stabilized cross-linked hydrogel of the presentinvention is sufficient to elicit a specific sequence of biologicalresponses, such as promoting tissue regeneration and vasculogenesis,without the addition of exogenous drugs or growth factors. In fact, thebioactive hydrogel matrix of the present invention can be substantiallyfree, even completely free, of exogenous drugs or growth factors whenused for vascularization or tissue regeneration. This intrinsicallybioactive hydrogel, as a result of its unique structure, provides a cellattachment scaffold that modulates subsequent cellular activity, such astissue regeneration and vasculogenesis.

The stabilized cross-linked hydrogel behaves similarly when used inother aspects of tissue regeneration. The hydrogel provides a stabilizedstructural lattice that facilitates cell retention and multiplication inareas with tissue damage. This is due in part to the intrinsicbioactivity of the hydrogel, which furthers the regenerative process.This is especially useful in applications where the success orfunctioning of an implanted medical device is dependent upon itsintegration with the surrounding tissue. The intrinsic bioactivity ofthe cross-linked hydrogel immobilized to the surface of the medicaldevice not only reduces incidence of rejection by the host resultingfrom inflammatory response, immune response, etc., but it also increaseshealing and tissue regeneration in the site surrounding the implanteddevice.

The immobilized bioactive hydrogel matrix surface coating utilized ineach of the embodiments described herein may be comprised solely of thetwo high molecular weight components. Preferably, each of theembodiments described herein incorporates additional components such asthe enhancing agents utilized in the preferred embodiments describedabove. Table 1 below lists preferred components present within theimmobilized bioactive hydrogel matrix surface coatings of the presentinvention along with suitable concentrations as well as preferredconcentrations for each component. Note that the concentrations listedin Table 1 for gelatin and dextran would also be suitable foralternative polyglycan and polypeptide components.

TABLE 1 Concentration Preferred Component Range Concentration L-glutamicacid 2 to 60 mM 15 mM L-lysine 0.5 to 30 mM 5 mM Arginine 1 to 40 mM 10mM Gelatin 0.01 to 40 mM 2 mM L-cysteine 5 to 500 μM 20 μM EDTA 0.01 to10 mM 4 mM Dextran 0.01 to 10 mM 0.1 mM

As noted above, the present invention provides numerous benefitsincluding eliciting vascularization at a localized site, modulatinglocalized wound healing response, and providing suitable means ofdeveloping a retrievable cell implantation device for cell-basedtherapeutics. Additional benefits may include the following: reducedscarring associated with degradation of bioerodible suture materials;improvement in the performance and long-term function of extravascularsensors such as glucose sensors routinely used for insulin deliverysystems; improvement in the rate of healing, durability, and mechanicalproperties around structural implants such as artificial joints andtendons; reduced pain and associated complications arising from postsurgical adhesions especially during abdominal or spinal injury; andimproved integration between natural tissues and implanted structures(i.e. teeth, porous hydroxyapatite or ceramic materials for bonerepair).

EXPERIMENTAL

The present invention is more fully illustrated by the followingexamples, which are set forth to illustrate the present invention andare not to be construed as limiting thereof.

Example 1

A beaker with an internal volume of 50 mL was equipped with two copperelectrodes at a 2.5 cm separation. The beaker was filled with an aqueoussolution of liquid thermoreversible hydrogel containing dextran andgelatin. A potential difference of 18 V was applied across the cell. Ahydrogel complex consisting of the covalently cross-linkedthermoreversible hydrogel formulation immediately formed on the surfaceof the anode, and the thickness of the film increased with increasingtime. The sterile hydrogel was insoluble in water at 37° C., wasadherent to the underlying substrate and conformed to the surface of theanodic metal.

One skilled in the art would readily recognize the utility of thismethod for producing adherent hydrogel coatings on metallic substratessuch as titanium meshes used for reconstructive surgery. Such bioactivehydrogel coatings are expected to improve the vascularity andosteointegration of the implant.

Example 2

Activated biomaterial surfaces suitable for having the hydrogel matrixcross-linked thereto can be prepared by copolymerization of monomerscontaining bifunctional groups, one of which is protected. For example,the monomer glycidyl methacrylate can be copolymerized using freeradical initiation with other acrylates to form hydrogels, and hydrogelfilms. Poly(2-hydroxyethyl methacrylate-co-glycidylmethacrylate)-poly(HEMA-GMA) hydrogel films can be prepared byUV-initiated photopolymerization with α,α′-azoisobutyronitrile (AIBN) asan initiator, preferably under an inert atmosphere at 25° C. The epoxidecontent of the hydrogel films can be varied by varying the relativeratio of HEMA to GMA. For example, films with a high density of epoxidescan be prepared by mixing 0.2 mL of HEMA, 0.8 mL GMA, 1 mL isopropylalcohol, 10 mg AIBN (as a polymerization initiator), and 3.0 mL of 0.1Mphosphate buffer (pH=7.0). The resulting mixture is stirred andequilibrated at 25° C. for 15 min in a thermostated water bath. Themixture can be then poured into the mold and exposed to long-waveultraviolet radiation for 20 min. After polymerization, poly(HEMA-GMA)films can be washed several times with distilled water and cut intocircular pieces with a biopsy punch. The functional epoxy group carryingpoly(HEMA-GMA) film disks (10 g wet weight, diameter=1.0 cm) formed asdescribed above are equilibrated in phosphate buffer (50 mM, pH=8.0) for2 hours, and transferred to a container holding the thermoreversiblehydrogel held at 39° C. Immobilization of the thermoreversible hydrogelto the surface of the biomaterial film can be carried out at 39° C. withfrequent agitation. The poly(HEMA-GMA) films coated with athermoreversible hydrogel can be removed and washed to removenon-covalently attached hydrogel materials.

Many modifications and other embodiments of the invention will come tomind to one skilled in the art to which this invention pertains havingthe benefit of the teachings presented in the foregoing descriptions andthe associated drawings. Therefore, it is to be understood that theinvention is not to be limited to the specific embodiments disclosedherein and that modifications and other embodiments are intended to beincluded within the scope of the appended claims. Although specificterms are employed herein, they are used in a generic and descriptivesense only and not for purposes of limitation.

1-20. (canceled)
 21. A coated implant comprising: an implant having asurface and being adapted for providing one or both of structural andmechanical support to contacted tissue; and a hydrogel matrix comprisinga polyglycan, a polypeptide, and at least one enhancing agent selectedfrom the group consisting of polar amino acids, divalent cationchelators, and combinations thereof; wherein the hydrogel matrix issolid or semi-solid at physiological temperature and physiological pH;and wherein the hydrogel matrix is overlying the surface of the implantand is immobilized thereon by at least one of the polyglycan andpolypeptide being covalently cross-linked to the implant surface. 22.The coated implant according to claim 21, wherein the polyglycan isdextran or oxidized dextran, and the polypeptide is gelatin.
 23. Thecoated implant according to claim 21, wherein the implant is formed of athermoplastic polymer.
 24. The coated implant according to claim 21,wherein the implant is a mesh.
 25. The coated implant according to claim21, wherein the polyglycan is a polysaccharide or a sulfatedpolysaccharide.
 26. The coated implant according to claim 25, whereinthe polysaccharide is selected from the group consisting of dextran,heparan, heparin, hyaluronic acid, alginate, agarose, carrageenan,amylopectin, amylose, glycogen, starch, cellulose, and chitin; andwherein the sulfated polysaccharide is selected from the groupconsisting of heparan sulfate, chondroitin sulfate, dextran sulfate,dermatan sulfate, and keratan sulfate.
 27. The coated implant accordingto claim 21, wherein the polypeptide is a tissue-derived polypeptideselected from the group consisting of collagens, gelatins, keratin,decorin, aggrecan, and glycoproteins.
 28. The coated implant accordingto claim 21, wherein the at least one enhancing agent comprises at leastone polar amino acid selected from the group consisting of tyrosine,cysteine, serine, threonine, asparagine, glutamine, aspartic acid,glutamic acid, arginine, lysine, histidine, and mixtures thereof. 29.The coated implant according to claim 21, wherein the at least oneenhancing agent comprises a divalent cation chelator selected from thegroup consisting of ethylenediaminetetraacetic acid and salts thereof.30. The coated implant according to claim 21, wherein the polyglycan andpolypeptide are covalently cross-linked to each other.
 31. A coatedtissue supporting mesh comprising: a mesh implant having a surface andbeing adapted for providing one or both of structural and mechanicalsupport to contacted tissue; and a hydrogel matrix comprising apolyglycan, a polypeptide, and at least one enhancing agent selectedfrom the group consisting of polar amino acids, divalent cationchelators, and combinations thereof; wherein the hydrogel matrix issolid or semi-solid at physiological temperature and physiological pH;and wherein the hydrogel matrix is overlying the surface of the meshimplant and is immobilized thereon by at least one of the polyglycan andpolypeptide being covalently cross-linked to the mesh surface.
 32. Thecoated tissue supporting mesh according to claim 31, wherein thepolyglycan is dextran or oxidized dextran, and the polypeptide isgelatin.
 33. The coated tissue supporting mesh according to claim 31,wherein the mesh is formed of a thermoplastic polymer.
 34. The coatedtissue supporting mesh according to claim 31, wherein the mesh is formedof a metallic material.
 35. The coated tissue supporting mesh accordingto claim 31, wherein the polyglycan is a polysaccharide or a sulfatedpolysaccharide.
 36. The coated tissue supporting mesh according to claim35, wherein the polysaccharide is selected from the group consisting ofdextran, heparan, heparin, hyaluronic acid, alginate, agarose,carrageenan, amylopectin, amylose, glycogen, starch, cellulose, andchitin; and wherein the sulfated polysaccharide is selected from thegroup consisting of heparan sulfate, chondroitin sulfate, dextransulfate, dermatan sulfate, and keratan sulfate.
 37. The coated tissuesupporting mesh according to claim 31, wherein the polypeptide is atissue-derived polypeptide selected from the group consisting ofcollagens, gelatins, keratin, decorin, aggrecan, and glycoproteins. 38.The coated tissue supporting mesh according to claim 31, wherein the atleast one enhancing agent comprises at least one polar amino acidselected from the group consisting of tyrosine, cysteine, serine,threonine, asparagine, glutamine, aspartic acid, glutamic acid,arginine, lysine, histidine, and mixtures thereof.
 39. The coated tissuesupporting mesh according to claim 31, wherein the at least oneenhancing agent comprises a divalent cation chelator selected from thegroup consisting of ethylenediaminetetraacetic acid and salts thereof.40. The coated tissue supporting mesh according to claim 31, wherein thepolyglycan and polypeptide are covalently cross-linked to each other.41. A coated implant comprising: a flexible implant comprising athermoplastic polymer and having a surface; and a hydrogel matrixcomprising a polyglycan, a polypeptide, and at least one enhancing agentselected from the group consisting of polar amino acids, divalent cationchelators, and combinations thereof; wherein the hydrogel matrix issolid or semi-solid at physiological temperature and physiological pH;and wherein the hydrogel matrix is overlying the surface of the flexibleimplant and is immobilized thereon.
 42. The coated implant according toclaim 41, wherein the polyglycan is dextran or oxidized dextran, and thepolypeptide is gelatin.
 43. The coated implant according to claim 41,wherein the implant is a mesh.
 44. The coated implant according to claim41, wherein the polyglycan is a polysaccharide or a sulfatedpolysaccharide.
 45. The coated implant according to claim 44, whereinthe polysaccharide is selected from the group consisting of dextran,heparan, heparin, hyaluronic acid, alginate, agarose, carrageenan,amylopectin, amylose, glycogen, starch, cellulose, and chitin; andwherein the sulfated polysaccharide is selected from the groupconsisting of heparan sulfate, chondroitin sulfate, dextran sulfate,dermatan sulfate, and keratan sulfate.
 46. The coated implant accordingto claim 41, wherein the polypeptide is a tissue-derived polypeptideselected from the group consisting of collagens, gelatins, keratin,decorin, aggrecan, and glycoproteins.
 47. The coated implant accordingto claim 41, wherein the at least one enhancing agent comprises at leastone polar amino acid selected from the group consisting of tyrosine,cysteine, serine, threonine, asparagine, glutamine, aspartic acid,glutamic acid, arginine, lysine, histidine, and mixtures thereof. 48.The coated implant according to claim 41, wherein the at least oneenhancing agent comprises a divalent cation chelator selected from thegroup consisting of ethylenediaminetetraacetic acid and salts thereof.49. The coated implant according to claim 41, wherein the polyglycan andpolypeptide are covalently cross-linked to each other.